Neck-worn physiological monitor

ABSTRACT

The invention provides a neck-worn sensor that is a single, body-worn system that measures the following parameters from an ambulatory patient: heart rate, pulse rate, pulse oximetry, respiratory rate, temperature, thoracic fluid levels, stroke volume, cardiac output, and a parameter sensitive to blood pressure called pulse transit time. From stroke volume, a first algorithm employing a linear model can estimate the patient&#39;s pulse pressure. And from pulse pressure and pulse transit time, a second algorithm, also employing a linear algorithm, can estimate systolic blood pressure and diastolic blood pressure. Thus, the sensor can measure all five vital signs along with hemodynamic parameters. It also includes a motion-detecting accelerometer, from which it can determine motion-related parameters such as posture, degree of motion, activity level, respiratory-induced heaving of the chest, and falls.

FIELD OF THE INVENTION

The invention relates to sensors that measure physiological signals frompatients and the use of such sensors.

BACKGROUND OF THE INVENTION

There are a number of physiological parameters that can be assessed bymeasuring physiological or physiologically influenced electrical signalsfrom a patient. Some signals, such as thoracic bioimpedance (TBI) andelectrocardiogram (ECG) waveforms, are measured with electrodes thatattach to the patient's skin. Processing of these waveforms yieldsparameters such as heart rate (HR), respiration rate (RR), heart ratevariability (HRV), stroke volume (SV), cardiac output (CO), andparameters related to thoracic fluids, e.g. thoracic fluid content(TFC). Certain physiological conditions can be identified from theseparameters when they are obtained at a single point in time; othersrequire assessment over some period of time to identify trends in theparameters. In both cases, it is important to obtain the parametersconsistently and with high repeatability and accuracy.

Some conditions require measuring parameters over a relatively short ormodest period of time. For example, Holter monitors can characterizevarious types of cardiac arrhythmias by measuring HR, HRV, and ECGwaveforms over a 24 to 48-hour period of time. On the other hand,chronic diseases such as congestive heart failure (CHF) and end-stagerenal disease (ESRD) can require periodic measurements of fluids andweight throughout the patient's life. Not surprisingly, patientcompliance typically decreases as the measurement period increases. Thisis particularly true when measurements are made outside of aconventional medical facility, e.g., at the patient's home or in aresidential facility such as a nursing home.

Measuring some physiological parameters does not require a high degreeof precision and/or consistency in the body location at which ameasurement is taken. For example, measuring a patient's temperature isfrequently performed with an oral thermometer that is simply placedsomewhere under the tongue. Here, the exact placement of the thermometerdoes not have a big impact on the measured temperature value. Likewise,parameters that depend on time-dependent features in waveforms, such asHR, which depends on the time-dependent variation of R-R intervals inthe ECG waveforms, are relatively insensitive to sensor positioning. Inthis case, the R-R intervals show almost no variation with positioningof electrodes on the patient's thoracic cavity. Blood pressure, incontrast, shows some sensitivity to measurement location. When measuredwith a sphygmomanometer, the blood pressure value is relativelyinsensitive to the general alignment of the cuff over the brachialartery, but will vary when measured at other locations on the body, suchas the wrist, thigh, or even the opposing arm. On the contrary,measuring amplitude-dependent features in waveforms, such as TFC, willbe strongly dependent on the positioning of electrodes. In this case,the value of TFC depends strongly on the impedance between theelectrodes, and this in turn will vary with the electrodes' placement.Here, deviation in day-to-day placement of electrodes can result inmeasurement errors, particularly when trends of the measured parametersare extracted. This, in turn, can lead to misinformation, nullify thevalue of such measurements, and thus negatively impact treatment.

Known Devices and Relevant Physiology

Medical devices that measure time-dependent ECG and TBI waveforms frompatients typically connect through cables or lead wires to disposableelectrodes adhered at various locations on a patient's body. Analogcircuits within a given device, which are typically located remote fromthe patient's body in the device, process the signals to generate thewaveforms. With further analysis, such waveforms yield parameters suchas HR, TFC, SV, CO, and RR. Other systems within a given medical devicemight measure vital signs such as pulse oximetry (SpO2), pulse rate(PR), systolic (SYS) and diastolic (DIA) blood pressure, and temperature(TEMP).

Disposable electrodes that measure ECG and TBI waveforms are typicallyworn on the patient's chest or legs and include: i) a conductivehydrogel that contacts the patient; ii) a Ag/AgCl-coated eyelet thatcontacts the hydrogel; iii) a conductive metal post that connects to alead wire or cable extending from the device; and iv) an adhesivebacking that adheres the electrode to the patient. Unfortunately, duringa measurement, the lead wires can pull on the electrodes if the deviceis moved relative to the patient's body, or if the patient ambulates andsnags the lead wires on various objects. Such pulling can beuncomfortable or even painful, particularly where the electrodes areattached to hirsute parts of the body, which can inhibit patientcompliance with long-term monitoring. Moreover, such pulling can degradeor even completely eliminate adhesion of the electrodes to the patient'sskin, thereby degrading or completely destroying the ability of theelectrodes to sense the physiolectrical signals at various electrodelocations.

Some devices that measure ECG and TBI waveforms are worn entirely on thepatient's body. These devices have been improved to feature simple,patch-type systems that include both analog and digital electronicsconnected directly to underlying electrodes. Such devices are typicallyprescribed for relatively short periods of time, e.g. for a time periodranging from a few days to several weeks. They are typically wireless,and include technologies such as Bluetooth® transceivers to transmitinformation over a short range to a second device, which then includes acellular radio to transmit the information to a web-based system.

Measurement of SpO2 values is almost always done from the patient'sfingers, earlobes, or in some cases toes. In these cases, patients wearan optical sensor to measure photoplethysmogram (PPG) waveforms, whichare then processed to yield SpO2 and PR. TEMP is typically measured witha thermometer inserted in the patient's mouth.

Assessing TFC, weight, and hydration status is important in thediagnosis and management of many diseases. For example, ESRD occurs whena patient's kidneys are no longer able to work at a level needed forday-to-day life. The disease is most commonly caused by diabetes andhigh blood pressure, and is characterized by swings in SYS and DIA alongwith a gradual increase in fluids throughout the body. Patientssuffering from ESRD typically require hemodialysis or ultrafiltration toremove excess fluids. Thus to characterize ESRD, accurate measurement ofthese TFC can eliminate the need for empirical clinical estimations thatoften lead to over-removal or under-removal of fluid during dialysis,thereby preventing hemodynamic instability and hypotensive episodes(Anand et al., “Monitoring Changes in Fluid Status With a WirelessMultisensor Monitor: Results From the Fluid Removal During AdherentRenal Monitoring (FARM) Study,” Congest Heart Fail. 2012; 18:32-36). Asimilar situation exists with respect to CHF, which is a complicateddisease typically monitored using a “constellation” of physiologicalfactors, i.e., fluid status (e.g. TFC), vital signs (i.e. HR, RR, TEMP,SYS, DIA, and SpO2), and hemodynamic parameters (e.g. CO, SV). Accuratemeasurement of these parameters can aid in managing patients,particularly for dispersing diuretic medications, thereby reducingexpensive hospital readmissions (Packer et al., “Utility of ImpedanceCardiography for the Identification of Short-Term Risk of ClinicalDecompensation in Stable Patients With Chronic Heart Failure,” J Am CollCardiol 2006; 47:2245-52).

CHF is a particular type of heart failure (HF), which is a chronicdisease driven by complex pathophysiology. In general terms, thiscondition occurs when SV and CO are insufficient in adequately perfusingthe kidneys and lungs. Causes of this disease are well known andtypically include coronary heart disease, diabetes, hypertension,obesity, smoking, and valvular heart disease. In systolic HF, ejectionfraction (EF) can be diminished (<50%), whereas in diastolic HF thisparameter is typically normal (>65%). The common signifyingcharacteristic of both forms of heart failure is time-dependentelevation of the pressure within the left atrium at the end of itscontraction cycle, or left ventricular end-diastolic pressure (LVEDP).Chronic elevation of LVEDP causes transudation of fluid from thepulmonary veins into the lungs, resulting in shortness of breath(dyspnea), rapid breathing (tachypnea), and fatigue with exertion due tothe mismatch of oxygen delivery and oxygen demand throughout the body.Thus, early compensatory mechanisms for HF that can be detected fairlyeasily include increased RR and HR.

As CO is compromised, the kidneys respond with decreased filtrationcapabilities, thus driving retention of sodium and water, and leading toan increase in intravascular volume. As the LVEDP rises, pulmonaryvenous congestion worsens. Body weight increases incrementally, andfluids may shift into the lower extremities. Medications for HF aredesigned to interrupt the kidneys' hormonal responses to diminishedperfusion, and they also work to help excrete excess sodium and waterfrom the body. However, an extremely delicate balance between these twobiological treatment modalities needs to be maintained, since anincrease in blood pressure (which relates to afterload) or fluidretention (which relates to preload), or a significant change in heartrate due to a tachyarrhythmia, can lead to decompensated HF.Unfortunately, this condition is often unresponsive to oral medications.In that situation, admission to a hospital is often necessary forintravenous diuretic therapy.

In medical centers, HF is typically detected using Doppler/ultrasound,which measures parameters such as SV, CO, and EF. In the homeenvironment, on the other hand, gradual weight gain measured with asimple weight scale is one method to indicate CHF. However, thisparameter is typically not sensitive enough to detect the early onset ofCHF—a particularly important time when the condition may be amelioratedby a change in medication or diet.

SV is the mathematical difference between left ventricular end-diastolicvolume (EDV) and end-systolic volume (ESV) and represents the volume ofblood ejected by the left ventricle with each heartbeat; a typical valueis about 70-100 mL. EF relates to EDV and ESV as described below inEquation 1:

$\begin{matrix}{{EF} = {\frac{SV}{EDV} = \frac{{EDV} - {ESV}}{EDV}}} & (1)\end{matrix}$

CO is the average, time-dependent volume of blood ejected from the leftventricle into the aorta and, informally, indicates how efficiently apatient's heart pumps blood through their arterial tree; a typical valueis about 5-7 L/min. CO is the product of HR and SV, i.e.,CO=SV×HR  (2)

CHF patients, and in particular those suffering from systolic HF, mayreceive implanted devices such as pacemakers and/or implantablecardioverter-defibrillators to increase EF and subsequent blood flowthroughout the body. These devices may include circuitry and algorithmsto measure the electrical impedance between different leads of thedevice. As thoracic fluid increases in the CHF patient, the impedancetypically is reduced. Thus, this parameter, when read by aninterrogating device placed outside the patient's body, can indicate theonset of heart failure.

Monitoring Solutions

As illustrated in FIG. 1, many of the above-mentioned parameters can beused as early markers of the onset of CHF. EF is typically low inpatients suffering from this chronic disease, and can be furtherdiminished by factors such as a change in physiology, an increase insodium in the patient's diet, or non-compliance with medications. Thisis manifested by a gradual decrease in SV, CO, and SYS that typicallyoccurs between two and three weeks before a hospitalization event. Thereduction in SV and CO diminishes perfusion to the kidneys. As describedabove, these organs then respond with a decrease in their filteringcapacity, thus causing the patient to retain sodium and water andleading to an increase in intravascular volume. This, in turn, leads tocongestion, which is manifested to some extent by a build-up of fluidsin the patient's thoracic cavity (e.g. TFC). Typically, a detectableincrease in TFC occurs about 1-2 weeks before hospitalization becomesnecessary. Body weight increases after this event, typically by betweenthree and five pounds, causing fluids to shift into the lowerextremities. At this point, the patient may experience an increase inboth HR and RR to increase perfusion. Nausea, dyspnea, and weight gaintypically grow more pronounced a few days before hospitalization becomesnecessary. As noted above, a characteristic of decompensated HF is thatit is often unresponsive to oral medications; thus, at this point,intravenous diuretic therapy in a hospital setting often becomesmandatory. A hospital stay for intravenous diuretic therapy typicallylasts about 4 days, after which the patient is discharged and the cycleshown in FIG. 1 can start over once more.

Not only is such cyclical pathology and treatment physically taxing onthe patient, it is economically taxing on society as well. In thisregard, CHF and ESRD affect, respectively, about 5.3 million and 3million Americans, resulting in annual healthcare costs estimated at $45billion for CHF and $35 billion for ESRD. CHF patients account forapproximately 43% of annual Medicare expenditures, which is more thanthe combined expenditures for all types of cancer. Somewhatdisconcertingly, roughly $17 billion of this is attributed to hospitalreadmissions. CHF is also the leading cause of mortality for patientswith ESRD, and this demographic costs Medicare nearly $90,000/patientannually. Thus, there understandably exists a profound financialincentive to keep patients suffering from these diseases out of thehospital. Starting in 2012, U.S. hospitals have been penalized forabove-normal readmission rates. Currently, the penalty has a cap of 1%of payments, growing to over 3% in the next three years.

Of some promise, however, is the fact that CHF-related hospitalreadmissions can be reduced when clinicians have access to detailedinformation that allows them to remotely titrate medications, monitordiet, and promote exercise. In fact, Medicare has estimated that 75% ofall patients with ESRD and/or CHF could potentially avoid hospitalreadmissions if treated by simple, effective programs.

Thus, with the aim of identifying precursors to conditions such as CHFand ESRD, physicians can prescribe monitoring solutions to patientsliving at home. Typically such solutions include multiple, standardmedical devices, e.g. blood pressure cuffs, weight scales, and pulseoximeters. In certain cases, patients use these devices daily and in asequential manner, i.e. one device at a time. The patient then calls acentral call center to relay their measured parameters. In more advancedsystems, the devices are still used in a sequential manner, butautomatically connect through a short-range wireless link (e.g.Bluetooth®) to a “hub”, which then forwards the information off to acall center. Often the hub features a simple user interface that posesbasic questions to the patient, e.g. questions concerning their diet,how they are feeling, and whether or not medications were taken.

Patients can also wear ambulatory cardiac monitors for periods of timeranging from several days to weeks to characterize cardiac conditions,such as arrhythmias. Such devices are called Holter or event monitors,and measure parameters such as HR, HRV, and ECG waveforms. Theytypically include a collection of chest-worn ECG electrodes (typically 3or 5); an ECG circuit that collects analog signals from the ECGelectrodes and converts them into multi-lead ECG waveforms; and aprocessing unit that analyzes the ECG waveforms to determine cardiacinformation. Typically, the patient wears the entire system on theirbody, but such systems can be awkward, cumbersome, and/or uncomfortable,e.g., due to the electrodes pulling the patient's skin/hair. Someambulatory systems may simply include on-board memory that storesinformation for retrieval at a later time, at which point theinformation is analyzed to generate a report describing the patient'scardiac rhythm. More modern systems include wireless capabilities totransmit ECG waveforms and other numerical data through a cellularinterface to an Internet-based system, where the report is generatedusing automated algorithms. In most cases, the report is imported intothe patient's electronic medical record (EMR), which avails the reportto cardiologists or other clinicians who can then use it to helpcharacterize and treat the patient.

In order for such monitoring to be therapeutically effective, however,it is important for the patient to use their equipment consistently,both in terms of the duration and manner in which it is used.Less-than-satisfactory consistency with the use of any medical device(in terms of duration and/or methodology) may be particularly likely inan environment such as the patient's home or a nursing home, wheredirect supervision may be less than optimal.

SUMMARY OF THE INVENTION

In view of the foregoing, it would be beneficial to improve patientcompliance with at-home monitoring. Here, compliance indicates theregularity and manner in which a patient uses a device. A sensoraccording to the invention, which facilitates monitoring a patient forHF, CHF, ESRD, cardiac arrhythmias, and other diseases in both home andclinical environments, could achieve this goal. The sensor is worn likea conventional necklace, comfortably around the neck, and features amechanical mechanism that ensures consistent placement when used on adaily basis, thereby improving the repeatability and reproducibility ofits measurements. Additionally, the sensor makes simultaneousmeasurements of multiple parameters, and thus obviates the need to usemultiple devices. Both of these features may improve patient compliance.

The sensor is able to detect the early onset of these and otherdiseases, thereby providing clinicians information that, when acted on,may prevent hospitalization. More particularly, the invention features aneck-worn sensor that is an integrated, body-worn system, which measuresthe following parameters from a patient: HR, PR, SpO2, RR, TEMP, athoracic fluid index (TFI), SV, CO, and a parameter sensitive to bloodpressure called pulse transit time (PTT). From SV, a first algorithmemploying a linear model can estimate the patient's pulse pressure (PP).And from PP and PTT, a second algorithm, also employing a linearalgorithm, can estimate SYS and DIA. Thus, the sensor, acting alone, canmeasure all five vital signs (HR/PR, SpO2, RR, TEMP, and SYS/DIA) alongwith hemodynamic parameters (SV, CO, TFI). Trends in some of theseparameters, e.g. SV, SYS, and TFI, may predict the onset of HF (e.g.CHF) before its severity is such that the patient requires admission toa hospital. By measuring this constellation of properties in thepatient's home and then wirelessly transmitting them to a clinician forevaluation, the sensor facilitates timely medical intervention that mayultimately keep the patient out of the hospital.

The sensor also includes a motion-detecting accelerometer, from which itcan determine motion-related parameters such as posture, degree ofmotion, activity level, respiratory-induced heaving of the chest, andfalls. The sensor can operate additional algorithms to process themotion-related parameters to measure vital signs and hemodynamicparameters when motion is minimized and below a pre-determinedthreshold, thereby reducing artifacts. Moreover, the sensor estimatesmotion-related parameters such as posture to improve the accuracy ofcalculations for vital signs and hemodynamic parameters.

Disposable electrodes attach directly to the sensor to secure it inclose proximity to the patient's body without bothersome cables. Inparticular, the electrodes are provided in patches, with each electrodepatch containing two electrode regions to measure ECG and TBI waveforms.The patches easily and releasably connect to circuit boards containedwithin the sensor by means of magnets that are electrically connected tothe circuit boards to provide signal-conducting electrical couplings.Prior to use, the electrodes are simply held near the circuit boards,and magnetic attraction causes the electrode patches to snap into properposition, thereby ensuring proper positioning of the electrodes on thepatient's body.

Using light-emitting diodes operating in the red (e.g. 600 nm) andinfrared (e.g. 800 nm) spectral regions, the sensor measures SpO2 and PRby pressing lightly against capillary beds in the patient's chest.Operating in a reflection-mode geometry, the sensor measures PPGwaveforms with both red and infrared wavelengths. SpO2 is processed fromalternating and static components of these waveforms. PR, in turn, canbe calculated from neighboring pulses, typically from the PPG waveformgenerated with infrared light, as this typically has a relatively highsignal-to-noise ratio.

All analog and digital electronics associated with these variousmeasurements are directly integrated into the sensor. This means that asingle, unobtrusive component—shaped like a piece of conventionaljewelry instead of a bulky medical device—measures a robust set ofparameters that can characterize a patient using both one-time andcontinuous measurements. Measurements can take place over just a fewminutes or several hours, and can be made in medical facilities and athome. The sensor includes a simple LED in its base (i.e. sensing)portion, which is located near the center of the chest when worn by thepatient. The sensor also includes a wireless transmitter (operatingBluetooth® and/or 802.11a/b/g/n) than sends data to, e.g., aconventional mobile device (e.g. cellular telephone, tablet computer,desktop/laptop computer, or plug-in hub).

The sensor measures all of the above-mentioned properties whilefeaturing a comfortable, easy-to-wear form factor that resembles a pieceof conventional jewelry. It is lightweight (about 100 grams) andbattery-powered. During use, it simply drapes around the neck, where thedisposable electrodes hold it in place, as described in more detailbelow. Flexible, conductive elements resembling strands in aconventional necklace power on the sensor, hold it in place, and alsoensure that it is consistently positioned when used on a daily basis.Moreover, the patient's neck is a location that is unobtrusive,comfortable, removed from the hands, and able to bear the weight of thesensor without being noticeable to the patient. The neck and thoraciccavity are also relatively free of motion compared to appendages such asthe hands and fingers, and thus a sensor affixed to the neck regionminimizes motion-related artifacts. Moreover, such artifacts arecompensated for, to some degree, by the accelerometer within the sensor.And because the sensor resembles jewelry (e.g., a necklace) and istherefore considerably less noticeable or obtrusive than variousprior-art devices, emotional discomfort over wearing a medical deviceover an extended period of time is reduced, thereby fostering long-termpatient compliance with a monitoring regimen.

The sensor's form factor is designed for comfort and ease of use, withthe ultimate goal of improving patient compliance so that theabove-mentioned parameters can be measured in a continuous manner and ona day-to-day basis. The system is targeted for elderly, at-homepatients, e.g. those suffering from chronic conditions such as HF, CHF,ESRD and related cardiac diseases, diseases of the kidneys, diabetes,and chronic obstructive pulmonary disease (COPD).

Thus, in one aspect, the invention features a sensor for simultaneouslymeasuring SYS, SV, and TFI from a patient. The sensor features a sensingportion having a flexible housing, and an elongated securement memberthat extends from the sensing portion to pass around the patient's neck.The elongated securement member has sufficient length to support thesensing portion generally against the sternal portion of the patient'schest when the sensor is in use and operating. It is configured toposition the sensing portion in a consistent location on the patient'sbody, thus optimizing the repeatability and reproducibility of eachmeasurement. As described above, this is particularly important whentrends are extracted from parameters measured on a daily basis.

The sensor features at least one pair of electrode contact pointsdisposed within the housing, with each pair of electrode contact pointscomprising a current-injecting electrode contact point and avoltage-sensing electrode contact point. Also disposed within thehousing is an analog ECG circuit in contact with at least one pair ofelectrode contact points and configured to generate an analog ECGwaveform based on a sensed voltage. An analog impedance circuit, alsodisposed within the housing and in electrical contact with at least onepair of electrode contact points, is configured to generate an analogimpedance waveform based on the sensed voltage. During a measurement,the waveforms pass to a digital processing system featuring amicroprocessor and analog-to-digital converter. This system digitizesthe analog ECG and TBI waveforms to generate, respectively, digital ECGand TBI waveforms.

Within the housing are systems for monitoring blood pressure (e.g. SYS),SV, and TFI. Each system features an algorithm that processes at leastone of the ECG and TBI waveforms to arrive at these values. For example,the blood pressure-monitoring system uses a first algorithm tocollectively process the digital TBI and ECG waveforms to determine avalue of SYS. The SV-monitoring system uses a second algorithm toprocess the digital TBI waveform to determine a value of SV. And theTFI-measuring system uses a third algorithm to process the digitalimpedance waveform to determine a value of TFI.

In embodiments, the elongated securement member features a battery,conductive wires for supplying voltage and ground to the sensingcomponent, and a clasp assembly at its distal end. During a measurement,the clasp assembly powers on the sensing component when it is attachedto the sensing component. Here, the sensing component includes a circuitthat prevents power from being supplied to the sensing component whenthe clasp assembly and sensing component are detached. The circuit,e.g., may be a control system that is configured to control the supplyof electrical power from a battery, whereby electrical power is suppliedto the analog and digital circuitry when the clasp assembly is mated tothe second end region of the housing. In embodiments, the clasp assemblyincludes a first connector, the sensing component includes a secondconnector, and the clasp assembly powers on the sensing component whenthe first connector connects to the second connector. The connectors,e.g., can be magnets. These can be positioned and arranged to releasablyhold the clasp assembly to the housing by way of magnetic attraction.Furthermore, an electrical circuit is completed when the magnets contacteach other, which allows electrical current to flow through the sensorsystem so as to power circuitry within the sensor to perform themeasurements described above.

In another aspect, the invention features a sensor for measuring trendsin SYS, SV, and TFI measured at different times by the sensing portion.Here, the elongated securement member positions the sensing portion inroughly the same location on the patient's body to improve therepeatability and reproducibility of repeated measurements. In anotheraspect, the invention features a sensor for generating alerts indicatingan onset of heart failure by measuring trends in SYS, SV, and TFI wheneach of these parameters trends to a lower value.

And in yet another aspect, the invention features a system for measuringphysiological signals from a patient that includes a flexible housing,configured to be worn around a patient's neck, that includes a firstcomponent enclosing a first circuit board, a second component enclosinga second circuit board, a third component enclosing a third circuitboard, a fourth component enclosing a first electrical conductor, and afifth component enclosing a second electrical conductor. A sensor withthis configuration is ideal for conforming to the contours present onthe thoracic region of some patients. At least one of the first, second,and third circuit boards features either an analog circuit configured tomeasure at least one time-dependent analog waveform from the patient, ora digital processing system featuring a microprocessor and ananalog-to-digital converter that digitizes at least one time-dependentanalog waveform to generate a time-dependent digital waveform. One ofthe circuit boards also includes a sensor for monitoring thephysiological signal from the patient. It operates an algorithm toprocess the time-dependent digital waveform to determine thephysiological signal.

In specific embodiments of the invention, an algorithm collectivelyprocesses the digital ECG and TBI waveforms to measure SYS from a pulsetransit time (PTT). Here, the algorithm: 1) processes the digital ECGwaveform to determine a first time point; 2) processes the digital TBIwaveform to determine a second time point; 3) analyzes the first andsecond time points to determine PTT; and 4) analyzes the PTT todetermine a value of SYS. In embodiments, the algorithm may also: 1)process SV to determine PP; and 2) process PTT and PP to determine SYSand DIA. An algorithm may also process the digital TBI waveform to: 1)extract an amplitude of a derivatized value of the TBI waveform's ACcomponent; 2) extract an amplitude of the TBI waveform's DC component;3) extract an estimated injection time; and 4) collectively process theamplitudes of the AC and DC components, along with the injection time,to determine SV. Furthermore, the amplitude of a DC component of thedigital TBI waveform may be processed to determine the value of TFI.

In alternate embodiments, the sensor can calculate a parameter calledvascular transit time (VTT) from fiducial points (e.g. the base ormaximum value of the derivatized waveform) of each pulsatile componentcontained in both the PPG and TBI waveforms. VTT can then be used inplace of PTT to calculate SYS. This approach has certain advantages,namely in that VTT lacks certain time components related to the cardiaccycle—called ‘systolic time intervals—that do not depend on bloodpressure. Inclusion of such components may thus add errors to the bloodpressure measurement. Examples of such systolic time intervals arepre-injection period (PEP) and the isovolumetric contraction time (ICT).

In still other embodiments, the sensor or the web-based system cancalculate a time-dependent change or deviation in any of theabove-mentioned parameters (e.g. SV, CO, SYS, DIA, TFI) to predict theonset of a disease, e.g. CHF.

In specific embodiments of sensors according to any of these aspects ofthe invention, flexibility is imparted to the sensing portion of thesensor, so that it can conform to the individual curvatures of differentpatients’ chests, by forming the housing in two or more segments, whichare connected to each other by flexible connector segments. Rigidcircuit boards located within the various housing segments are connectedto each other via flexible circuits, which pass through the flexibleconnector segments.

BRIEF DESCRIPTION OF THE DRAWINGS

Embodiments of the invention will now be described in greater detail inconjunction with the figures, in which:

FIG. 1 is a timeline illustrating how detectable physiologicalparameters precede and can be used to predict the onset of CHF;

FIG. 2 is a schematic diagram illustrating the use of a neck-worn sensoraccording to the invention as part of a system for remote monitoring ofpatients;

FIG. 3A is a photograph of an embodiment of a neck-worn sensor accordingto the invention, as illustrated in use in FIG. 2, with a standard blackenclosure;

FIG. 3B is a photograph of an embodiment of a neck-worn sensor accordingto the invention, as illustrated in use in FIG. 2, with a decorativefloral enclosure;

FIG. 4 is a photograph of an embodiment of a neck-worn sensor accordingto the invention as illustrated in use in FIG. 2, with a transparenthousing to reveal its internal components;

FIG. 5 is a perspective view of the neck-worn sensor shown in FIGS. 3and 4, with its housing removed to show arrangement of its internalcircuit boards and other electrical components;

FIG. 6A is a photograph of the sensor shown in FIGS. 3 and 4, with itsmagnetic clasp portion attached to a battery charger;

FIGS. 6B and 6C are photographs of the sensor shown in FIGS. 3 and 4,with, respectively, its magnetic clasp attached and detached from itsbase sensing portion;

FIG. 7 is a photograph of the battery charger, configured to charge abattery housed within the sensor when mated with the magnetic claspportion shown in FIGS. 6A, 6B, and 6C;

FIG. 8 is a schematic circuit diagram of a circuit that controlspowering of the sensor and charging of its battery;

FIG. 9 is a photograph of the sensor in use, with a transparent housingto reveal internal components thereof, and which front view is providedfor reference for FIG. 9A;

FIG. 9A is a perspective view of pulse oximetry circuitry included inthe circled portion of FIG. 9;

FIG. 10 is a photograph of the sensor in use, with a transparent housingto reveal internal components thereof and which front view is providedfor reference for FIG. 10A;

FIG. 10A is a photograph of the sensor's backside showing the use ofmagnetically attached electrode patches;

FIGS. 11A-11E are, respectively, graphs illustrating the followingtime-dependent waveforms: FIG. 11A) ECG waveform; FIG. 11B) DCcomponents of the TBI waveform; FIG. 11C) unfiltered AC components ofthe TBI waveform; FIG. 11D) low pass-filtered AC components of the TBIwaveform; and FIG. 11E) band pass-filtered AC components of the TBIwaveform;

FIGS. 12A and 12B are, respectively, graphs illustrating the followingtime-dependent waveforms: FIG. 12A) ECG and derivatized TBI waveforms;and FIG. 12B) low-passed filtered AC components of the TBI waveform;

FIG. 13 is a schematic diagram of a circuit within the sensor used tomeasure TBI waveforms;

FIG. 14 is a corresponding flow chart illustrating how the sensormeasures TBI waveforms and calculates values of SV in the presence ofmotion;

FIG. 15 is a graph showing the correlation of TFI measured by the sensorand a related, impedance-derived parameter measured with a referencedevice, the BioZ;

FIG. 16 is a graph showing the correlation of SV measured by the sensorand SV measured with a reference device, the BioZ; and,

FIG. 17 is a graph showing the correlation of RR measured by the sensorand RR measured with a reference device, a vital sign monitor with anend-tidal CO2 sensor.

DETAILED DESCRIPTION

Remote Patient-Monitoring System

As illustrated in FIG. 2, a sensor 10 according to the invention is onecomponent of an overall system 1 designed to facilitate remotemonitoring of—and, if required, therapeutic intervention on behalf of—apatient 12. As will be explained in greater detail below, the sensor 10measures numerical and waveform data, and then sends this wirelessly toa gateway system 18, and from there to a web-based system 26 that can beviewed at a remote facility 14 such as a hospital, medical clinic,nursing facility, and/or eldercare facility. There, clinicians canevaluate or further process the data to evaluate the patient 12.

In particular, after measuring and processing the patient'sphysiological parameters as addressed below, the sensor 10 automaticallytransmits data through an internal Bluetooth® wireless transmitter(identified schematically at 16) to a gateway system 18, which suitablymay be either a 2Net™ system 20 available from Qualcomm, Inc. or atablet computer or other handheld device 22 running a customizedgraphical user interface (GUI). Either type of gateway 18 (i.e., a 2Net™system 20 or a tablet computer/handheld device 22) preferably runs adownloadable software application that accesses the gateway's internalBluetooth® driver. The 2Net™ system 20 lacks any display, but features aseries of LEDs that indicate when it receives and then transmitsinformation for further review. The GUI on the tablet computer/handhelddevice 22 renders data on the device's screen, and also guides a patientthrough the measurement process. Both gateway devices use internal WiFiand/or cellular transmission channels 24 to transmit data to an IPaddress associated with the web-based system 26, which is typicallyoperated at the remote facility 14. The web-based system 26 displaysdata from multiple patients for clinicians, and may also include aninterface for individual patients. For example, the web-based system 26may display ECG and TBI waveforms; trended numerical data; and thepatient's medical history and demographic information. A clinicianviewing the web-based system 26 may, for example, analyze the data andthen call the patient 12 to have him or her adjust his/her medicationsand/or diet.

Alternatively, the sensor 10 can automatically transmit data viaBluetooth® transmitter to a personal computer (not illustrated), whichthen uses a wired or wireless Internet connection to transmit data tothe web-based system 26. In such a configuration, the personal computerruns a software program configured to download data from the sensor 10;to display the data for the patient 12 in an easy-to-understand format;and to forward the data to the computer server that runs the web-basedsystem 26 for relatively complex analysis as described below. In yetanother configuration, the sensor 10 can connect to the personalcomputer via a USB connection, and download and forward data to theweb-based system 26 using a wired Ethernet connection, as describedabove.

Basic Construction of Sensor

FIGS. 3-10 show embodiments of a sensor 10 according to the inventiondesigned to be worn by a patient 12. The sensor 10 is typically wornaround the patient's neck 28 so that it rests against their sternum,similar to a necklace or other neck-adorning jewelry. The sensor 10features a sensing portion 30 and a securement member 32 (or securementmembers in an alternate embodiment, not illustrated). As illustrated,the securement member 32 extends from a first end 34 of the sensingportion 30 and attaches to a second end 36 of the sensing portion 30.The securement member 32 is long enough to pass behind the patient'sneck 28 and to hold the sensing portion 30 in proper position forsensing electrodes attached to its rear, patient-facing surface to beattached to the proper locations on the patient's chest. This ensuresthat the sensing portion 30 is placed in approximately the same positionfor each measurement made on a particular patient, and that it is heldin proper position to acquire the relevant bioelectric signals, asexplained more fully below. Additionally, the securement member 32houses a battery in battery compartment 38, which is positionedgenerally in the middle of the securement member 32 (lengthwisespeaking) such that it is positioned inconspicuously behind thepatient's neck 28 when the sensor 10 is worn.

In other, non-illustrated embodiments, the securement member could besplit in the middle, with flexible yet shape-retaining “branches”extending from the first and second ends 34, 36 of the sensing portion30 so as to pass behind the patient's neck 28, but not connect, muchlike a physician's stethoscope. In that case, the battery compartmentcould be located in one of the branches or, alternatively, in thesensing portion 30 of the sensor 10. In still further non-illustratedembodiments, a securement member might not be included, in which caseattachment of the electrodes to the patient's body would, by itself, beused to hold the sensor in position. Ultimately, however, where asecurement member is provided to facilitate positioning of the sensingportion 30 on the patients' body, what is important is simply that thesecurement member should be configured to pass at least substantiallyaround the patient's neck 28 (which includes a configuration in whichlateral halves of the securement member pass posteriorly over thetrapezius muscles without curving medially toward the spine). In otherwords, the securement member 32 passes sufficiently over the trapeziusregion and/or behind the neck to support the sensing portion 30 andprevent it from falling before the sensing portion 30 is secured to thepatient's body via the electrodes, as described more fully below.

The sensing portion 30 is constructed in two or more sections orsegments, e.g. a central segment 42 and two outboard segments 40 a and40 b, to the rear of which electrode patches are attached as describedbelow. The segments are connected to each other by means of flexibleconductor segments 56 a, 56 b, which in turn are encased in flexiblehousing 46 and 48. The flexible conductor segments 56 a, 56 b aretypically made from a polymeric material, e.g. Kapton® flexible printedcircuits available from the DuPont Corporation. Such materials areessentially a flexible, polymeric film that encases one or more thinconducting members, which are typically made from copper. Each of thesegments 40 a, 40 b, and 42 includes, respectively, a rigid circuitboard 52 a, 52 b, and 54 (best shown in FIG. 5) populated with discreteelectrical circuit components, described in more detail below. The rigidcircuit boards 52 a, 52 b, and 54 connect to one another via theflexible conductor segments 56 a, 56 b, which each include 20 conductivemembers.

The rigid circuit boards 52 a, 52 b, and 54 are each encased inside of arigid protective housing segments 53 a, 53 b, 55, and the flexibleconductor segments 56 a, 56 b are encased within the flexible connectorsegments 46 and 48. The material of the protective housing segments 53a, 53 b, and 55 is shown as clear or transparent in FIGS. 4, 6A-6C, 9,and 10 and 10A in order to better show the underlying circuitcomponents. The protective housing segments 53 a, 53 b, and 55 are moretypically made from opaque plastic, as illustrated in FIGS. 3A and 3B,which contributes to the overall aesthetically pleasing appearance ofthe sensor 10. In other embodiments, as shown particularly in FIG. 3B,the protective housing segments can be made from plastic coated with adecorative pattern, such as a floral pattern. Or in other embodiments,clip-on coverings or decals can be applied to the protective housingsegments to occasionally change the appearance of the sensor, e.g. tomake it match a particular article of clothing. This can also make thesensor appear like conventional jewelry, which in turn may make apatient more likely to wear it.

Suitably, the connector segments 46 and 48, which may be formed asrubber boots designed to snap into respectively opposing ends of theprotective housing segments 52 a, 52 b, and 54, are typically made fromsoft, flexible material such as silicone rubber. Generally speaking,such a configuration of the sensing portion 30 serves to hold thesensing electrodes at their proper positions before they are adhered tothe patient's chest, while allowing the sensing portion 30 to conform tothe different curvatures of the physiological region upon which itrests.

As further illustrated in FIGS. 3 and 4, a transparent or translucentplastic window 57 located on the top, anteriorly facing surface ofcentral housing segment 55 covers an underlying LED 59, which serves asa simple user interface for the patient 12. For example, the LED 59 canradiate different colors of the visible spectrum, and blink them atdifferent frequencies, to indicate when the sensor 10 is turned on,making a measurement, charging, running on low power, completed with ameasurement, etc.

As shown most clearly in FIGS. 5 and 6A-6C, a pair of conductors 58(e.g., a twisted pair of braided-strand or solid-core wiring) suppliespower and ground from the sensor's battery to circuitry housed withinthe segments 40 a, 40 b, and 42 of the sensing portion 30. Theconductors 58 also supply a control (“enable”) signal, described in moredetail with respect to FIG. 8, that serves to power on the sensor. Morespecifically, the conductors 58 are electrically connected to and extendfrom electrical components on one of the rigid circuit boards (e.g.,circuit board 52 a). The conductors 58, which are sheathed in siliconeor other soft material that will be comfortable when worn around thepatient's neck, are also electrically connected at their distal ends tocomponents on the clasp circuit board 64 of a magnetic-switch claspassembly 62.

The magnetic-switch clasp assembly 62 houses the clasp circuit board 64and a magnetic connector 68. Suitably, the silhouette or planformcontours of the clasp assembly 62 match those on the upper surface ofthe protective housing segment 53 b encasing the rigid circuit board.Furthermore, that protective housing segment has a slight recess 66 inits upper or anterior-facing surface and into which the clasp assembly62 fits. The recess 66 allows the clasp assembly 62 to mate with theprotective housing segment 53 b in a way that forms a mostly smooth,generally continuous upper surface, as best shown in FIGS. 3A, 3B, and6C.

As further shown in FIGS. 5 and 6B-6C, a first clasp magnet 68 isattached to the clasp circuit board 64 within the clasp assembly 62 andis partially exposed. A second clasp magnet 70 is attached to the rigidcircuit board 52 b within the outboard sensor segment 40 b and ispartially exposed/accessible through the housing 53 b. Suitably, thefirst and second clasp magnets 68, 70 are oppositely polarizedrare-earth magnets that are coated with a film of conductive metal(e.g., chromium). They are electrically connected to circuitry locatedon the clasp circuit board 62 and the rigid circuit board 52 b,respectively.

The first and second clasp magnets 68 and 70 are positioned on the claspcircuit board 64 and rigid circuit board 52 b, respectively, so thatwhen the clasp assembly 62 is mated with the protective housing segment53 b as shown, for example, in FIG. 6C, the magnets 68 and 70 arebrought into engagement with each other. Furthermore, the first andsecond clasp magnets 68, 70 are oriented so that opposite polaritiesthereof will face each other when the clasp assembly 62 is mated withthe protective housing segment 53 b. As a result, the first and secondclasp magnets 68 and 70 are able to releasably secure themagnetic-switch clasp assembly 62 to the outboard sensor segment 40 b.

So engaging the magnetic-switch clasp assembly 62 to the sensor segment40 b—and, more particularly, connecting the clasp magnets 68 and 70 toeach other—completes a circuit (described below in conjunction with FIG.8) that functions as a control circuit within the clasp circuit board64. This powers on the sensor 10 and drives the control circuit tosupply power from the battery to the sensor portion 30, thereby causingthe sensor 10 to turn on and initiate a measurement, as described inmore detail below.

From this description of the conductors 58 and the magnetic-switch claspassembly 62 located at the end of them, it will be appreciated that thepair of conductors 58 function as the securement member 32 by means ofwhich the sensor 10 is worn around the patient's neck 28 in theillustrated embodiment. Furthermore, as shown most clearly in FIGS. 4,5, and 6A and as alluded to above, the battery compartment 38 is locatedgenerally halfway along the length of the securement member32/conductors 58. Suitably, the battery compartment can be formed as acylindrical capsule into which a rechargeable lithium ion battery fits.(For the disclosed embodiment, the battery preferably generates 4.2Vwhen it is fully charged and 3.3V when it is depleted.) Although notspecifically illustrated, the capsule may be formed from twohalf-cylindrical portions—possibly joined along lengthwise-extendingedges that form a hinge—that snap together to enclose the battery withinthe compartment 38. In other non-illustrated embodiments, the batterycan be removable, and be replaced in the compartment. Electricalcontacts (not illustrated) are located at opposite ends of the batterycompartment 38 and engage the positive and negative terminals of thebattery when the battery is inserted into the battery compartment 38.This configuration allows a completed circuit to be formed when themagnetic-switch clasp assembly 62 is engaged with the sensor segment 40b as described above.

Advantageously with this configuration, a battery charger 74 asillustrated in FIGS. 6A, 7, and 8 can be provided with the sensor 10 tocharge the sensor's lithium ion battery. The battery charger 74 has aplastic receptacle member 76 with a depression-shaped receptacle region78 configured to engage with the exterior-facing surface of the claspassembly 62. A charger magnet 80 is attached to or partially embeddedwithin the receptacle region 78 and is positioned within the receptacleregion so as to engage with the switch clasp magnet 68 when the claspassembly 62 is brought into engagement with the receptacle member 76.Furthermore, the charger magnet 80 is oriented such that oppositepolarities of the charger magnet 80 and the clasp magnet 68 will faceeach other then the clasp assembly 64 engages with the receptacle member76. Like the clasp magnets 68, 70, the charger magnet 80 is a rare-earthmagnet that is coated with a film of conductive metal (e.g., chromium).

A signal-conducting cable 81 passes partially through and out ofreceptacle member 76 and terminates in a standard USB plug 84, with oneelectrical conductor (not shown) that is contained within the cable 81being electrically connected to the charger magnet 80. Furthermore, apair of spring-loaded contact pins 82 (commonly called ‘pogo pins’)extend into the receptacle region 78, and are positioned to make contactwith a corresponding pair of electrical contacts 83 that are located onthe clasp circuit board 64 and that are accessible via a correspondingpair of holes formed in the clasp assembly 62, as shown in FIG. 6B. Twoadditional electrical conductors (not shown) that are contained withinthe cable 81 are electrically connected to the pins 82 and supply power(+5V) and ground to the electrical contacts 83 on the clasp circuitboard 64 when the magnetic-switch clasp assembly is properly mated withthe battery charger 74. The USB plug 84 can plug into the USB outlet ofa personal computer or a common AC/DC converter 93 (with a female USBsocket) that converts, e.g., 120V AC current from a standard outlet to+5V DC.

To charge the sensor's battery, the anterior-facing surface of the claspassembly 62 is brought into engagement with the receptacle region 78,thus causing the clasp magnets 68 and 80 to engage each other andreleasably and magnetically secure the clasp assembly 62 within thereceptacle 78. This action causes a circuit to be completed and causesthe control circuit in the clasp circuit board to power on the sensor,as explained immediately below. When this happens, DC power from theAC/DC converter 93 or USB outlet of a personal computer passes throughthe USB plug 84, along the conducting cable 81, through the pins 82, andto the electrical contacts 83 disposed on the clasp circuit board 64.This action fully charges the battery over a period of about 4 hours.

Operation and one possible arrangement of the control circuit 86 areillustrated in FIG. 8 (note that the circuit shown in the figure is asimplified version of the actual circuit schematic corresponding to thesensor). The circuit enables powering of the sensor using the claspassembly 62, as well as charging of the battery with the battery charger74. In both cases, voltage directly from the battery (typically between3.7 and 4.2V) is connected directly to the WAKE UP line in the conductor88. To power the sensor, the magnetic-switch clasp assembly 62 engageswith the sensor segment 40 b, causing clasp magnets 68 and 70 to connectto each other. This supplies voltage directly from the WAKE UP line toENABLE pins in both the voltage regulator 89 and microprocessor 87, thusenabling these components to function as designed. Once this occurs, thevoltage regulator 89 trims an input voltage from the battery to 3.3V,which leaves the voltage regulator from the OUT line to power themicroprocessor 87 and most other digital components in the sensor. Atthis point, the microprocessor is enabled, and can carry out allnecessary functions, such as switching on circuitry and blinking anyLEDs, even if the clasp assembly 62 is disconnected. A similar situationexists with the battery charger 74 when the clasp assembly 62 is broughtinto engagement with the receptacle region 78, thus causing the magnets68 and 80 to contact each other. When this occurs, the microprocessor ispowered and enabled as described above, and in response shuts down allpower-consuming analog circuitry in the sensor. This allows the batteryto charge in an efficient manner when it is placed in the batterycharger 74, as described above.

As shown in FIGS. 9 and 9A, the sensor 10 can also include a standardpulse oximetry circuit 100 such as the one described in U.S. Pat. No.8,437,824, the contents of which are incorporated by reference in theirentirety. Using a non-invasive, optical measurement, the pulse oximetrycircuit 100 generates a value of SpO2. It may be located on the rear,patient-facing side of one of the sensor's outboard circuit boards,e.g., on circuit board 52 a, and generally near the attachment pointsfor a pair of electrodes that are used with the sensor (as explained ingreater detail below). The pulse oximetry circuit 100 drives red andinfrared LEDs in an alternating, pulsatile manner and controls alight-sensitive, photodetector diode, as generally known in the art.(FIG. 9A illustrates an integrated module, in which the LEDs andphotodetector are not independently discernible; because this technologyis well known in the field, they are not identified individually in thefigures.)

The pulse oximetry circuit 100 is configured to operate in a reflectionmode, meaning that the LEDs and light-sensitive diode are positioned soas to receive radiation from the same direction. It measures PPGwaveforms from capillary beds in the patient's chest to generate a valueof SpO2, as is described in more detail below. This is in contrast toconventional pulse oximetry sensors in which the LEDs and thelight-sensitive diode are positioned across from each other, with aspace into which fits a body part (e.g., a finger or an earlobe) beinglocated between the LEDs and the light-sensitive diode. Thus, the pulseoximetry circuit detects and measures radiation emitted by the diodesthat has been reflected off of capillary beds (i.e., in the chest)before arriving at the light-sensitive diode.

Because the pulse oximetry sensor is incorporated into the overallsensor 10, the pulse oximetry optical sensor can connect comfortably tothe patient's chest to measure SpO2 values in an effective manner thateliminates “cable clutter” and frees the patient's hands and fingers(where pulse oximetry measurements typically are taken) for otherpurposes. An additional benefit of this configuration is reduction ofmotion artifacts, which can distort PPG waveforms and cause erroneousvalues of SpO2 to be reported. This reduction of motion artifacts is dueto the fact that during everyday activities, the chest typically movesless than the hands and fingers, and subsequent artifact reductionultimately improves the accuracy of SpO2 values measured from thepatient.

Finally with regard to the basic construction of a sensor 10 accordingto the invention, both a three-axis digital accelerometer and atemperature sensor (not specifically identified) are provided on thecentral circuit board 54 to measure, respectively, three time-dependentmotion waveforms (along x, y, and z-axes) and TEMP values.

Bioelectric Signal-Acquiring and Signal-Processing Components

As noted above, a sensor 10 according to the invention monitors apatient by means of bioelectrical signals. In particular, electricalcircuitry on the rigid circuit boards 52 a, 52 b, and 54 as well as theflexible circuits 56—working in conjunction with disposable electrodesthat are attached to the outboard rigid circuit boards 52 a and 52 b asdescribed more fully below—are designed to measure impedance and, fromthat impedance, yield the TBI waveform. The ECG waveform is measured ina known manner, and other relevant parameters can then be calculatedfrom the TBI waveform as addressed below.

Thus, as shown in FIGS. 10 and 10A, each of the outboard rigid circuitboards 52 a and 52 b has a pair of circular, electrode-retainingrare-earth magnets 114 attached to it. (A two-part electrode patch 116that magnetically connects to them hides the electrode-retaining magnetson circuit board 52 b). The electrode-retaining magnets 114 arepositioned within similarly sized circular openings in the rear walls ofthe protective housing segments 53 a and 53 b, and thus they areaccessible from the rear exterior of the outboard segments 40 a and 40b. Furthermore, each of the electrode-retaining magnets 114 has aconductive metal coating (e.g. chromium) and is electrically connectedto the circuitry located on its respective circuit board 52 a or 52 b.As a result, the magnets 114 are able to conduct electrical signalssensed by the electrodes into the TBI and ECG analog circuits of thesensor 10.

Two-part electrode patches 116, which are suitable for use with a sensor10 according to the invention, are generally known in the art.Construction of similar electrode patches is described, for example, incurrently U.S. Patent Application No. 61/747,864, filed Dec. 31, 2013(entitled “Body-worn Sensor for Characterizing Patients with HeartFailure”), the contents of which are incorporated herein by reference.In general, in such an electrode patch 116, an adhesive backing supportseach of two distinct and electrically uncoupled conductive electrodes,and each of the electrodes has a sticky, conductive hydrogel thatcontacts the patient's skin. As shown in FIG. 10A, a small hole oraperture 130 on the order of 5 millimeters in diameter is formed throughthe electrode patch 116 to permit radiation associated with the pulseoximetry circuit 100 and the pulse oximetry measurement to pass throughthe electrode patch so that SpO2 values can be measured from capillarybeds in the patient's chest. For each electrode, the hydrogel contactsan eyelet, which is coated on one side with a thin layer of Ag/AgCl, andis connected on its other side to a metal post made from conductivestainless steel.

Because the stainless steel metal posts of the electrode patches 116 areattracted to the magnets 114, electrode patches 116 can be attached tothe sensor 10 relatively simply by holding an electrode patch near, andgenerally aligned with, one of the outboard sensor segments 40 a or 40b, with the electrodes near the magnets 114. Magnetic attraction betweenthe magnets 114 and the stainless steel posts will then cause theelectrode patch 116 to snap into place (but removably so), therebyeliminating the need for cumbersome snaps and rivets that can bedifficult for elderly patients to manipulate. Moreover, forcesassociated with pressing the electrodes' metal posts into thecorresponding electrode holders of known, prior art-devices can causecircuit components to pop off, thus impeding performance of the device.Both these conditions are ameliorated to some extent by using magnets.

Thus, with a sensor 10 constructed as per the invention, electrodepatches are connected to the sensing portion 30 first, and then they areconnected to the patient by pressing the sensing portion 30, and hencethe electrodes, against the patient's skin. Because the shape of thesensing portion 30 is essentially fixed (allowing for slight deviationas the sensing portion flexes at the flexible connector segments 46 and48), applying the electrodes to the patient's skin in this mannerfacilities relatively consistent and accurate location of the electrodeson the patient, thereby enhancing the accuracy—especially over time—andhence the value of the information obtained with the sensor 10.Additionally, proper location and spacing of the electrodes ensures bothTBI and ECG waveforms are acquired with high signal-to-noise ratios;this, in turn, leads to measurements that are relatively easy to analyzeand that, as a result, have optimized accuracy.

The circuit boards 52 a, 52 b, and 54 have arranged thereon a firstelectrical circuit for making an impedance-based measurement of TBIwaveforms that yield CO, SV, RR, and fluid levels, and a secondelectrical circuit for making differential voltage measurements of ECGwaveforms that yield HR and arrhythmia information. The first electricalcircuit 134 is illustrated schematically in FIG. 14; understanding ofthe arrangement and functionality of the components that comprise thefirst electrical circuit 134 will, however, be facilitated by apreceding explanation of the bioelectrical signal-acquisition andprocessing functions performed by the circuitry. Furthermore, the secondelectrical circuit is of a type that is well known in this particularart, and thus it is not described in significant detail herein.

As indicated schematically in FIGS. 10 and 10A, each electrode patch 116includes a current-injecting electrode region I1, I2 described in detailbelow. (It should be noted that just a single electrode patch 116 isshown in FIG. 10A but that during use of a sensor 10 according to theinvention, an electrode patch 116 is attached to the electrode-retainingmagnets 114 on both lateral sides of the sensor 10.) Each electrodepatch 116 also includes another, voltage-sensing electrode regions S1,S2. During a measurement, the S1, S2 regions sense bioelectric signals.These pass through a first circuit, which includes a collection offilters, amplifiers, and rectifying components that generate atime-dependent, analog TBI waveform that relates by Ohm's Law toimpedance encountered by the injected current. This is explained indetail below with respect to FIG. 13. Likewise, the bioelectric signalsalso pass through a second circuit, which features a differentialamplifier and filtering components that generate a time-dependent,analog ECG waveform. Such a circuit is standard in this particular art,and is thus not described in detail here.

Thus, to use the sensor 10 to monitor a patient's physiology, anelectrode patch 116 is first attached to the sensing portion 30 of thesensor 10, and then the electrodes are attached to the patient's chestas noted above. Ideally, each electrode patch 116 attaches just belowthe patient's collarbone, near the patient's left and right arms. Duringa measurement, the impedance circuit (described more fully below)injects a high-frequency, low-amperage current (I) into the patient'sskin through electrode regions I1, I2. Typically, the modulationfrequency is about 70 kHz, and the current is about 6 mA. The currentinjected by each electrode region I1, I2 is out of phase by 180° withrespect to the other. It encounters static (i.e. time-independent)resistance from components such as fluids and, to a lesser extent, bone,skin, and other tissue in the patient's chest. Additionally, blood andother fluids in the chest conduct the current to some extent. Bloodejected from the left ventricle of the heart into the aorta provides adynamic (i.e. time-dependent) resistance. As the largest artery passingblood out of the heart, the aorta has a dominant impact on the dynamicresistance; other vessels such as the superior vena cava, on the otherhand, will contribute to the dynamic resistance in a minimal way.

Electrode regions S1, S2 measure a time-dependent voltage (V) thatvaries with resistance (R) encountered by the injected current (I)according to Ohm's Law, shown below in Equation 3:V=I×R  (3)

During a measurement, the time-dependent voltage is filtered by theimpedance circuit and ultimately measured with an analog-to-digitalconverter within the electronic circuitry. This digitized voltage isthen processed to calculate SV using an equation such as that shownbelow in Equation 4, which is the Sramek-Bernstein equation, or amathematical variation thereof. Historically, parameters extracted fromTBI signals are fed into the equation, shown below, which is based on avolumetric expansion model taken from the aortic artery:

$\begin{matrix}{{SV} = {\delta\frac{L^{3}}{4.25}\frac{\left( {{{dZ}(t)}/{dt}} \right)_{\max}}{Z_{0}}{LVET}}} & (4)\end{matrix}$

In Equation 4, Z(t) represents the TBI waveform, δ representscompensation for body mass index, Z₀ is the base impedance, L isestimated from the distance separating the current-injecting andvoltage-measuring electrodes on the thoracic cavity, and LVET is theleft ventricular ejection time, which is the time separating the openingand closing of the aortic valve. An averaged value of Z₀ is equivalentto TFI and will vary with fluid levels. Typically, a high resistance(e.g. one above about 30Ω) indicates a dry, dehydrated state. In thatcase, the lack of conducting thoracic fluids increases resistivity inthe patient's chest. Conversely, a low resistance (e.g. one below about19Ω) indicates the patient has more thoracic fluids, and is possiblyoverhydrated. In that case, the abundance of conducting thoracic fluidsdecreases resistivity in the patient's chest. LVET can be determinedfrom the TBI waveform, or from the HR using an equation called“Weissler's Regression,” shown below in Equation 5, that estimates LVETfrom HR:LVET=−0.0017×HR+0.413  (5)

Weissler's Regression allows LVET to be estimated from HR determinedfrom the ECG waveform. This equation and several mathematicalderivatives, along with the parameters shown in Equation 4, aredescribed in detail in the following reference, the contents of whichare incorporated herein by reference: “Impedance Cardiography, Pulsatileblood flow and the biophysical and electrodynamic basis for the strokevolume equations,’ Bernstein, Journal of Electrical Bioimpedance, Vol.1, p. 2-17, 2010. Both the Sramek-Bernstein Equation and an earlierderivative of this, called the Kubicek Equation, feature a “staticcomponent,” Z₀, and a “dynamic component,” Z(t), the derivative of whichrelates to LVET and a (dZ/dt)_(max)/Z₀ value. These equations assumethat (dZ(t)/dt)_(max)/Z₀ represents a radial velocity of blood (withunits of Ω/s) due to volume expansion of the aorta.

Furthermore, during a measurement, the second electrical circuitmeasures an analog ECG waveform that is received by an internalanalog-to-digital converter within the onboard microprocessor 136 (shownin FIG. 13). An algorithm processes the waveform to determine HR andSYS/DIA, as described in more detail below. Additionally, themicroprocessor 136 analyzes this signal simply to determine whether theelectrode patches are properly adhered to the patient and that thesystem is operating satisfactorily. An ECG waveform of high quality(e.g. having a signal-to-noise ratio>5) indicates that electrode patchesare properly adhered, while one of low quality (e.g. having asignal-to-noise ratio<5) indicates the opposite. Once this state isachieved, the first and second electrical circuits generatetime-dependent analog waveforms that a high-resolution analog-to-digitalconverter within the electronics module receives and sequentiallydigitizes to generate time-dependent digital waveforms. Typically, thesewaveforms are digitized with 16-bit resolution over a range of about10V. The microprocessor receives the digital waveforms and processesthem with computational algorithms, written in embedded computer codesuch as C or Java, to generate values of CO, SV, TFI, and HR, as isdescribed in more detail below

Further still, during a measurement, the light-sensitive diode in thepulse oximetry circuit 100 receives radiation from the associated LEDsthat reflects off of tissue. Signals from the light-sensitive diode passthrough amplifier and filter circuitry to yield PPG waveforms emanatingfrom the red and infrared radiation. These waveforms are digitized withan analog-to-digital converter and then processed to extract fiducialpoints, as described in the above-referenced-and-incorporated '824patent. The fiducial points are then processed with an algorithm thatimplements Equation 6, below, to determine a SpO2 value.

$\begin{matrix}{R = \frac{{{red}({AC})}/{{red}({DC})}}{{{infrared}({AC})}/{{infrared}({DC})}}} & (6)\end{matrix}$

In Equation 6, the red (AC) and red (DC) represent, respectively,parameters extracted from the AC and DC components of the PPG waveformmeasured with the red LED. A similar case holds for the infrared (AC)and infrared (DC) values. The term “AC signals”, as used herein, refersto a portion of a PPG waveform that varies relatively rapidly with time,e.g., the portion of the signal caused by pulsations in the patient'sblood. “DC signals,” in contrast, are portions of the PPG that arerelatively invariant with time, e.g., the portion of the signaloriginating from scattering off of components such as bone, skin, andnon-pulsating components of the patient's blood.

More specifically, AC signals are measured from a heartbeat-inducedpulse present in both waveforms. The pulse represents a pressure wave,launched by the heart, which propagates through the patient'svasculature and which causes a time-dependent increase in volume in botharteries and capillaries. When the pressure pulse reaches vasculatureirradiated by the oximeter's optical system, a temporary volumetricincrease results in a relatively large optical absorption according tothe Beer-Lambert Law. DC signals originate from radiation scatteringfrom static components such as bone, skin, and relatively non-pulsatilecomponents of both arterial and venous blood. Typically, only about 0.5%to about 1% of the total signal measured by the pulse oximetryphotodetector is attributable to the AC signal, with the remainder beingattributable to the DC signal. Separation of AC and DC signals istypically done with both analog and digital filtering techniques thatare well-known in the art.

The R value in Equation 6, which is sometimes called a “ratio of ratios”(RoR), represents a ratio of hemoglobin Hb to oxygenated hemoglobinHbO2. It equates an actual SpO2 value, which ranges from 0-100% O2, toan empirical relationship that resembles a non-linear equation. Aboveabout 70% O2, this equation typically yields values that are accurate towithin a few percentage points. On the other hand, while not necessarilyaccurate, measurements below this value still indicate a hypoxic patientin need of medical attention. Additional details for this calculationare described in the above-referenced-and-incorporated patent.

As noted above, the pulse oximetry circuit 100 operates in a reflectionmode. Therefore, radiation from the pulse oximetry diodes passes throughthe hole 130 in the electrode patch to tissue in the chest and thenreflects back before arriving at the pulse oximetry light-sensitivediode, where this component and the pulse oximetry circuit process it toform the requisite PPG waveforms needed for Equation 6. A conventionalPPG waveform measured with the above-described optical sensor features asequence of heartbeat-induced pulses, with the time duration separatingthe pulses being inversely related to PR. The heartbeat-induced pulsesrepresent blood pulsing in an underlying artery that absorbs (orreflects) incident radiation from the red and infrared LEDs. The PPGwaveform also includes a slowly varying baseline that is due tounderlying optical absorption by the blood. PPG waveforms emanating fromboth waveforms look similar, with that from infrared radiation typicallyhaving a relatively high signal-to-noise ratio.

FIGS. 11A-11E show ECG and TBI waveforms measured using the sensor 10.Waveforms shown in the figures are digital waveforms, sampled at a rateof 250 Hz. A typical ECG waveform (FIG. 11A) features a collection of“QRS complexes,” with each QRS complex representing an individualheartbeat; the waveform shown in FIG. 11A is similar to a standard ECGwaveform measured by a conventional vital sign monitor. FIGS. 11B-11Eshow different versions of the TBI waveform generated using analog anddigital filtering processes that are implemented by the sensor 10.Analog filtering is typically performed with a standard RC circuitlocated directly on the sensor's circuit boards, whereas digitalfiltering is typically performed using software algorithms operating onthe sensor's onboard microprocessor (shown in FIG. 13, which isdescribed below).

In general, analog and digital filtering techniques are well known inthe art. As applied in the context of the present invention, they can beused to isolate AC (i.e., rapidly varying) and DC (i.e., slowly varying)components of the waveforms, which, in turn, yield differentphysiological parameters. For example, the DC component of the TBIwaveform is sensitive to thoracic fluid levels and thus yields TFI. Inparticular, as thoracic fluid levels increase, the DC impedance measuredby the sensor 10, shown as a relatively flat line in FIG. 11B, willdecrease. The unprocessed AC component of the TBI waveform, shown inFIG. 11C, shows oscillating, time-dependent features corresponding todifferent physiological events (e.g. HR, RR) and that typically areisolated with digital filtering. As shown in FIG. 11D, processing the ACcomponent of the TBI waveform with a low-pass (LP) filter removesrelatively high-frequency components, thus yielding a relativelylow-frequency undulation attributable to RR. This waveform can befurther processed to identify periods of apnea, which are indicated bythe lack of undulations in the waveform after about 100 seconds.Similarly, processing the AC component of the TBI waveform with aband-pass (BP) filter, as illustrated in FIG. 11E, selectively removesany low-frequency undulations. This leaves periodic, heartbeat-inducedpulses attributable to blood flowing from the left ventricle into theaorta. These pulses can be processed to yield both SV and CO, asdescribed further below.

FIGS. 12A and 12B, in turn, show derivatized TBI and ECG waveformsmeasured with the sensor 10, as plotted over a short time window (i.e.,about 5 seconds) in FIG. 12A and a longer time window (i.e., about 60seconds) in FIG. 12B. As shown more clearly in FIG. 12A, individualheartbeats produce time-dependent pulses in both the ECG and TBIwaveforms. As is clear from the data, pulses in the ECG waveform precedethose in the TBI waveform. The ECG pulses, which each feature a sharp,rapidly rising QRS complex, indicate initial electrical activity incontractions in the patient's heart and, informally, the beginning ofthe cardiac cycle. The QRS complex is the peak of the ECG waveform. TBIpulses follow the QRS complex by about 100 ms and indicate blood flowthrough arteries in the patient's thoracic cavity. These signals aredominated by contributions from the aorta, which is the largest arteryin this region of the body. During a heartbeat, blood flows from thepatient's left ventricle into the aorta, and the volume of blood is theSV. Blood flow enlarges this vessel, which is typically very flexible,and also temporarily aligns blood cells from their normally randomorientation. Both of these mechanisms—enlargement of the aorta andtemporary alignment of the blood cells—improve electrical conductionnear the aorta, thus decreasing the electrical impedance as measuredwith TBI. The TBI waveform shown in FIG. 12A is the first mathematicalderivative of the raw TBI waveform; therefore, its peak represents thepoint of maximum impedance change.

A variety of time-dependent parameters can be extracted from the ECG andTBI waveforms. For example, as indicated in the upper portion of thefigure, it is well known that HR can be determined from the timeseparating neighboring ECG QRS complexes. Likewise, LVET can be measureddirectly from the TBI pulse. LVET is measured from the onset of thederivatized pulse to the first positive-going zero crossing. Alsomeasured from the derivatized TBI pulse is (dZ/dt)_(max), which is aparameter that is used to calculate SV as shown in Equation 4 above andas described in more detail in the above-cited-and-incorporatedliterature reference.

The time difference between the ECG QRS complex and the peak of thederivatized TBI waveform represents a PTT. This value can be calculatedfrom other fiducial points, particularly on the TBI waveform (such asthe base or midway point of the heartbeat-induced pulse). Typically,however, the peak of the derivatized waveform is used, as it isrelatively easy to develop a software beat-picking algorithm that findsthis fiducial point.

In other, non-illustrated embodiments, the sensor can calculate VTT fromfiducial points (e.g. the base or maximum value of the derivatizedwaveform) of each pulsatile component contained in both the PPG and TBIwaveforms. VTT can then be used in place of PTT to calculate SYS. Asdescribed above, this approach has certain advantages, namely in thatVTT lacks certain systolic time intervals (e.g. PEP, ICT) that do notdepend on blood pressure, and may thus add errors to the blood pressuremeasurement. In other words, SYS calculated from VTT may be moreaccurate than SYS calculated from PTT.

PTT (and VTT) correlates inversely to SYS and DIA, as shown below inEquations 7 and 8, where m_(SYS) and m_(DIA) are patient-specific slopesfor, respectively, SYS and DIA, and SYS_(cal) and DIA_(cal) are values,respectively, of SYS and DIA measured during a calibration measurement.(Without calibration, PTT only indicates relative changes in SYS andDIA.) A calibration can be provided with conventional means such as anoscillometric blood pressure cuff or in-dwelling arterial line. Thecalibration yields both the patient's immediate values of SYS and DIA.Multiple values of PTT and blood pressure can be collected and analyzedto determine patient-specific slopes m_(SYS) and m_(DIA), which relatechanges in PTT with changes in SYS and DIA. The patient-specific slopescan also be determined by using pre-determined values from a clinicalstudy and then combining these measurements with biometric parameters(e.g. age, gender, height, weight) collected during the clinical study.

$\begin{matrix}{{SBP} = {\frac{m_{SBP}}{PTT} + {SBP}_{cal}}} & (7) \\{{DBP} = {\frac{m_{DBP}}{PTT} + {DBP}_{cal}}} & (8)\end{matrix}$

In embodiments of a sensor 10 according to the invention, waveforms likethose shown in FIG. 12A are processed to determine PTT, which is thenused to determine either SYS or DIA according to Equations 7 and 8.Typically, PTT and SYS correlate to each other better than PTT and DIA,and thus this parameter is first determined. Then, PP is estimated fromSV, calculation of which is described below. Most preferably, instantvalues of PP and SV are determined, respectively, from the bloodpressure calibration and from the TBI waveform.

PP can be estimated from either the absolute value of SV, SV modified byanother property (e.g. LVET), or the change in SV. In the first method,a simple linear model is used to process SV (or, alternatively, SV×LVET)and convert it into PP. The model uses the instant values of PP and SV,determined as described above from a calibration measurement, along witha slope that relates PP and SV (or SV×LVET). The slope can be estimatedfrom a universal model that, in turn, is determined using a populationstudy. Alternatively, a slope tailored to the individual patient isused. Such a slope can be selected, for example, using biometricparameters describing the patient, as described above. Here, PP/SVslopes corresponding to such biometric parameters are determined from alarge population study and then stored in computer memory on the sensor10. When a particular sensor unit is assigned to a patient, theirparticular biometric data is entered into the system, e.g. using amobile telephone that transmits the data to the microprocessor in thesensor via Bluetooth® protocols. Then, an algorithm on the sensorprocesses the data and selects a patient-specific slope. Calculation ofPP from SV is described in the following reference, the contents ofwhich are incorporated herein by reference: “Pressure-Flow Studies inMan. An Evaluation of the Duration of the Phases of Systole,” Harley etal., Journal of Clinical Investigation, Vol. 48, p. 895-905, 1969. Asexplained in this reference, the relationship between PP and SV for agiven patient typically has a correlation coefficient (r) that isgreater than 0.9, which indicates excellent agreement between these twoproperties. Similarly, in the above-mentioned reference, SV is shown tocorrelate with the product of PP and LVET, with most patients showing anr value of greater than 0.93 and the pooled correlation value (i.e. thatfor all subjects) being 0.77. This last result indicates that a singlelinear relationship between PP, SV, and LVET may hold for all patients.

More preferably, PP is determined from SV using relative changes inthese values. Typically, the relationship between the change in SV andchange in PP is relatively constant across all subjects. Thus, similarto the case for PP, SV, and LVET, a single, linear relationship can beused to relate changes in SV and changes in PP. Such a relationship isdescribed in the following reference, the contents of which areincorporated herein by reference: “Pulse pressure variation and strokevolume variation during increased intra-abdominal pressure: anexperimental study,” Didier et al., Critical Care, Vol. 15:R33, p. 1-9,2011. Here, the relationship between PP variation and SV variation for67 subjects displayed a linear correlation of r=0.93, i.e., an extremelyhigh value for pooled results that indicates a single, linearrelationship may hold for all patients.

From such a relationship, PP is determined from the TBI-based SVmeasurement, and SYS is determined from PTT. DIA is then calculated fromSYS and PP.

The sensor 10 determines RR from both the TBI waveform and a motionwaveform generated by the accelerometer (called the ACC waveform). FIG.12B illustrates how the TBI waveform yields RR. In this case, thepatient's respiratory effort moves air in and out of the lungs, thuschanging the capacitance (and hence impedance) in the thoracic cavity.This time-dependent change maps onto the TBI waveform, typically in theform of oscillations or pulses that occur at a much lower frequency thanthe heartbeat-induced cardiac pulses shown in FIG. 12A. Thus, simplesignal-processing (e.g. filtering, beat-picking) of the low-frequency,breathing-induced pulses in the waveform yields RR.

Similarly, the ACC waveform will reflect breathing-induced movements inthe patient's chest. This results in pulses within the waveform thathave a similar morphology to those shown in FIG. 12B for the TBIwaveform. Such pulses can be processed as described above to estimateRR. RR determined from the ACC waveform can be used by itself, or it canbe processed collectively with RR as determined from the TBI waveform(e.g., using adaptive filtering) to improve accuracy. Such an approachis described in U.S. patent application Ser. No. 12/559,426 filed Sep.14, 2009 (entitled “Body-Worn Monitor For Measuring Respiration Rate”and published on Mar. 17, 2011 as U.S. Pub. 2011/0066062), the contentsof which are incorporated herein by reference. Furthermore, as shown inFIG. 12B, the baseline of the TBI waveform, called Z₀ or alternativelyTFI, can be easily determined. Z₀ is used to determine SV, as describedabove in Equation 4.

With the foregoing description of how the sensor 10 operates(computationally speaking) in mind, the circuitry used to generate,sense, and analyze the corresponding signals is easier to understand.Thus, FIG. 13 illustrates an analog circuit 134 that performs impedancemeasurement according to the invention. The figure shows just onepossible embodiment of the circuit 134; similar monitoring results canbe achieved using a design and collection of electrical components thatdiffer from those shown in the figure.

The circuit 134 has a first current-injecting electrode region 138 athat injects a high-frequency, low-amperage current (I1) into thepatient's thoracic cavity. This serves as the current source. Typically,a current pump 140 provides the modulated current, with the modulationfrequency typically being between 50 KHz and 100 KHz and the currentmagnitude being between 0.1 mA and 10 mA. Preferably, the current pump140 provides current with a magnitude of 4 mA that is modulated at 70kHz through the first current-injecting electrode region 138 a. A secondcurrent-injecting electrode region 138 b also injects a high-frequency,low-amperage current (I2) into the patient's thoracic cavity, with thecurrent injected by the second current-injecting region 138 b being 180°out of phase with respect to the current (I1) injected by the firstcurrent-injecting electrode region.

Sensing-electrode regions 142 a and 142 b sense the time-dependentvoltages encountered by the propagating current I1 and I2, respectively.These sensing-electrode regions are indicated in the figure as S1 andS2. Per Ohm's law as indicated above, the voltage sensed by thesesensing-electrode regions 142 a, 142 b divided by the magnitude of theinjected current yields a time-dependent resistance (i.e., impedance)that relates to blood flow in the aortic artery. As shown by thewaveform 144 in the figure, the time-dependent resistance features aslowly varying DC offset, characterized by Z₀, that indicates thebaseline impedance encountered by the injected current; for TBI, thiswill depend, for example, on the amount of thoracic fluids, along withthe fat, bone, muscle, and blood volume in the chest of a given patient.Z₀, which typically has a value between about 10 and about 150Ω, is alsoinfluenced by low-frequency, time-dependent processes such asrespiration. Such processes affect the inherent capacitance near thechest region that TBI measures and are manifested in the waveform bylow-frequency undulations, such as those shown in the waveform 144. Arelatively small (typically about 0.1-0.5Ω) AC component, Z(t) lies ontop of Z₀ and is attributed to changes in resistance caused by theheartbeat-induced blood that propagates in the brachial artery asdescribed in detail above. Z(t) is processed with a high-pass filter toform a TBI signal that features a collection of individual pulses 146that are ultimately processed to determine SV and CO as described above.

Voltage signals measured by the first sensing-electrode region 142 a(S1) and the second sensing-electrode region 142 b (S2) feed into adifferential amplifier 148 to form a single, differential voltage signalwhich is modulated according to the modulation frequency (e.g., 70 kHz)of the current pump 140. From there, the signal flows to a demodulator150, which also receives a carrier frequency from the current pump 140to selectively extract signal components that only correspond to the TBImeasurement. The collective function of the differential amplifier 148and the demodulator 150 can be accomplished using many differentcircuits designed to extract weak signals—such as the TBI signal—fromnoise. For example, these components can be combined to form somethingequivalent to a “lock-in amplifier” that selectively amplifies signalcomponents occurring at a well-defined carrier frequency. Or the signaland carrier frequencies can be deconvoluted in much the same manner asthat used in a conventional AM radio using a circuit featuring one ormore diodes. The phase of the demodulated signal may also be adjustedwith a phase-adjusting component 152 during the amplification process.In one embodiment of a sensor according to the invention, the ADS1298family of chipsets marketed by Texas Instruments may be used for thisapplication. This chipset features fully integrated analog front endsfor both ECG and impedance pneumography. The latter measurement isperformed with components for digital differential amplification,demodulation, and phase adjustment—such as those used for the TBImeasurement—that are integrated directly into the chipset.

Once the TBI signal is extracted, it flows to a series of analog filters154, 156, 158 within the circuit 134 that remove extraneous noise fromthe Z₀ and Z(t) signals. The first low-pass filter 154 (30 Hz) removesany high-frequency noise components (e.g. power line components at 60Hz) that may corrupt the signal. “Part” of the signal that passesthrough the filter 154, which represents Z₀, is ported directly to achannel in an analog-to-digital converter 160. The “remaining” part ofthe signal feeds into a high-pass filter 156 (0.1 Hz), which lets passhigh-frequency signal components responsible for the shape of individualTBI pulses 146. This signal then passes through a final low-pass filter158 (10 Hz) to further remove any high-frequency noise. Finally, thefiltered signal passes through a programmable gain amplifier (PGA) 162,which, using a 1.65V reference, amplifies the resultant signal with acomputer-controlled gain. The amplified signal represents Z(t) and isported to a separate channel of the analog-to-digital converter 160,where it is digitized alongside of Z₀. The analog-to-digital converterand PGA are integrated directly into the ADS1298 chipset describedabove. The chipset can simultaneously digitize waveforms such as Z₀ andZ(t) with 24-bit resolution and sampling rates (e.g. 500 Hz) that aresuitable for physiological waveforms. Thus, in theory, this one chipsetcan perform the function of the differential amplifier 148, demodulator150, PGA 162, and analog-to-digital converter 160. Reliance on just asingle chipset to perform these multiple functions ultimately reducesboth size and power consumption of the TBI circuit 134.

The microprocessor 136 receives digitized Z₀ and Z(t) signals through aconventional digital interface such as an SPI or I2C interface.Algorithms for converting the waveforms into actual measurements of SVand CO are performed by the microprocessor 136. The microprocessor 136also receives digital motion-related waveforms from the on-boardaccelerometer and processes these waveforms to determine parameters suchas the degree/magnitude of motion, frequency of motion, posture, andactivity level.

A corresponding flow chart of an algorithm 166 that functions usingcompiled computer code operating, e.g., on the onboard microprocessor136 is shown in FIG. 14. The algorithm 166 is used to measure TBIwaveforms and, from them, CO and SV in the presence of motion. Thecompiled computer code is loaded in memory associated with themicroprocessor and is run each time a TBI measurement is converted intonumerical values for CO and SV. The microprocessor typically runs anembedded real-time operating system, and the compiled computer code istypically written in a language such as C, C++, Java, or assemblylanguage. Each step S170-S185 in the algorithm 166 is typically carriedout by a function or calculation included in the compiled computer code.

Algorithms similar to that shown in FIG. 14 can be used to calculateother physiological parameters in the presence of motion such as SpO2,RR, HR, PR, PTT, and SYS/DIA based on PTT.

As for accuracy/reliability of the parameters measured and calculated inthis manner, FIGS. 15-17 show correlation plots indicating the efficacyof the sensor's measurements of TFI (FIG. 15), SV (FIG. 16), and RR(FIG. 17). In each plot, the value measured by the sensor is shown onthe y-axis, and the value measured by a reference device is shown on thex-axis. Perfect correlation is indicated by the line with a slope of m=1running through the plot. Each data point in the plots represents datafrom a unique patient, as measured during a clinical trial. For theplots showing TFI and SV, the reference is the Cardiodynamics BioZ,which is a device that uses impedance cardiography to measurecorresponding values of Z₀. For the plot showing RR, the referencedevice is a vital sign monitor performing a measurement of end-tidalCO2. As is clear from the figures, the sensor's measurements of itsrespective parameters correlate strongly with the reference techniques.

Finally, as shown in FIG. 13, both numerical and waveform data processedwith the microprocessor 136 are ported to a wireless transmitter 16 suchas a transmitter based on protocols like Bluetooth® or 802.11a/b/g/n.From there, the transmitter 16 sends data to an external receiver suchas a conventional cellular telephone, tablet, wireless hub (such asQualcomm's 2Net™ system), or personal computer. Devices like these canserve as a “hub” to forward data to an Internet-connected remote serverlocated, e.g., in a hospital, medical clinic, nursing facility, oreldercare facility, and there a clinician is able to evaluate the datafor early markers of CHF, ESRD, or other physiological conditions aswell as physical conditions (e.g., fallen, ambulating, etc.), asexplained above.

Other embodiments are deemed to be within the scope of the invention.For example, algorithms can process other waveforms, such as the PPG andECG waveforms, to extract parameters such as RR. In that case, thelow-frequency envelope of the waveform indicates RR. In otherembodiments, the reflective pulse oximetry system that measures SpO2 canbe replaced with an ear-worn optical sensor that connects to the sensorthrough a cable. Here, the sensor uses either reflective ortransmission-mode optical configurations to measure both the red andinfrared PPG waveforms. In other embodiments, algorithms that operate oneither the sensor or the gateway monitor trends in physiologicalparameters to determine the onset of a particular disease, e.g. CHF. Instill other embodiments, the electronics and mechanical componentswithin the sensor can be relocated within the sensor's geometry. Forexample, they can be moved from the back portion of the sensor to a sideportion proximal to the front of the patient's neck. Still further, thecontrol circuit could be relocated from the clasp assembly to thesensing portion of the sensor.

In other embodiments, the system shown above can take the form of a‘patch’ that directly adheres to a portion of a patient's body, asopposed to a ‘necklace’ that drapes around the patient's neck. The patchwould be similar in form to the necklace's base, although it may take onother shapes and form factors. It would include all the same sensors(e.g. sensors for measuring ECG, TBI, and PPG waveforms) and computingsystems (e.g. microprocessors operating algorithms for processing thesewaveforms to determine parameters such as HR, HRV, RR, BP, SpO2, TEMP,CO, SV, fluids) as the base of the necklace. However unlike the systemdescribed above, the battery to power the patch would be located in orproximal to the base, as opposed to the strands in the case of thenecklace. Also, in embodiments, the patch would include a mechanism suchas a button or tab functioning as an on/off switch. Alternatively, thepatch would power on when sensors therein (e.g. ECG or temperaturesensors) detect that it is attached to a patient.

In typical embodiments, the patch includes a reusable electronics module(shaped, e.g., like the base of the necklace) that snaps into adisposable component that includes electrodes similar to those describedabove. The patch may also include openings for optical and temperaturesensors as described above. In embodiments, for example, the disposablecomponent can be a single disposable component that receives thereusable electronics module. In other embodiments, the reusableelectronics module can include a reusable electrode (made, e.g., from aconductive fabric or elastomer), and the disposable component can be asimple adhesive component that adheres the reusable electrode to thepatient.

In preferred embodiments the patch is worn on the chest, and thusincludes both rigid and flexible circuitry, as described above. In otherembodiments, the patch only includes rigid circuitry and is designed tofit on other portions of the patient's body that is more flat (e.g. theshoulder).

Still other embodiments are within the scope of the invention.

What is claimed is:
 1. A sensor for measuring from a patient a systolicblood pressure (SYS), a diastolic blood pressure (DIA), and a pulseoximetry (SpO2), the sensor configured to be located on the patient'schest and comprising: a sensing portion having a flexible housingconfigured to be located on the patient's chest and enclosing a battery,wireless transmitter, and all of the sensor's sensing and electroniccomponents, including: at least two pairs of electrode contact pointsdisposed on a bottom surface of the flexible housing, with each pair ofelectrode contact points comprising a current-injecting electrodecontact point and a voltage-sensing electrode contact point; an analogelectrocardiogram (ECG) circuit contained entirely within the flexiblehousing and in electrical contact with a first voltage-sensing electrodecontact point from a first pair of electrode contact points and a secondvoltage-sensing electrode contact point from a second pair of electrodecontact points, the analog ECG circuit configured to generate an analogECG waveform based on sensed voltage; an analog impedance circuitcontained entirely within the flexible housing and in electrical contactwith current-injecting and voltage-sensing electrode contact points inboth the first and second pair of electrode contact points, the analogimpedance circuit being configured to generate an analog impedancewaveform based on sensed voltage; an optical system located on a bottomsurface of the flexible housing, the optical system comprising a lightsource configured to generate radiation in both the red and infraredspectral ranges that separately irradiates a portion of the patient'schest disposed underneath the flexible housing, and a photodetectorconfigured to detect radiation in the red spectral range that reflectsoff the portion of the patient's chest to generate a first analogphotoplethysmogram waveform (red-PPG), and to detect radiation in theinfrared spectral range that reflects off the portion of the patient'schest to generate a second analog photoplethysmogram waveform(infrared-PPG), with the optical system located between a firstelectrode contact point and a second electrode contact point of one ofthe two pairs of electrode contact points and configured to bereleasably connected to the patient's chest by the electrode contactpoints, such that the optical system irradiates the portion of thepatient's chest between the first electrode contact point and the secondelectrode contact point of one of the two pairs of electrode contactpoints; a digital processing system contained entirely within thehousing and comprising a microprocessor and an analog-to-digitalconverter, the digital processing system being configured to: 1)digitize the analog ECG waveform to generate a digital ECG waveform, 2)digitize the analog impedance waveform to generate a digital impedancewaveform, 3) digitize the analog red-PPG waveform to generate a digitalred-PPG waveform, and 4) digitize the analog infrared-PPG waveform togenerate a digital infrared-PPG waveform; a first sensor for measuringstroke volume (SV) disposed within the flexible housing, the firstsensor comprising a first processor configured to determine SV from amaximum value of a mathematical derivative of the digital impedancewaveform and a baseline value of the digital impedance waveform; asecond sensor for measuring blood pressure contained entirely within theflexible housing, the second sensor comprising a second processorconfigured to collectively process the digital impedance waveform andone of the digital red-PPG and infrared-PPG waveforms to determine: 1) afirst time point from one of the digital red-PPG and digitalinfrared-PPG waveforms, 2) a second time point from the digitalimpedance waveform, 3) a vascular transit time (VTT) from the temporaldifference between the first and second time points, 4) a value of SYSfrom a linear equation that includes an inverse value of VTT, 5) a valueof pulse pressure (PP) from the product of SV and a calibrationparameter relating this to PP, and 6) a value of DIA from themathematical difference between SYS and PP; and a third sensor formeasuring SpO2 contained entirely within the flexible housing, the thirdsensor comprising a third processor configured to determine a value ofSpO2 from alternating and static components of both the digital red-PPGwaveform and digital infrared-PPG waveform.
 2. The sensor of claim 1,wherein the flexible housing comprises two or more rigid housingsegments that are connected to each other by means of one or moreflexible connectors.
 3. The sensor of claim 2, wherein the analog ECGand analog impedance circuits and the digital processing system arelocated on rigid circuit boards disposed within the housing segments andthe analog ECG circuit, analog impedance circuit, and digital processingsystem are interconnected via one or more flexible conductors locatedwithin the one or more flexible connectors.
 4. The sensor of claim 3,wherein each of the one or more flexible connectors comprises a flexiblecircuit.
 5. The sensor of claim 4, wherein the analog ECG circuit andthe digital processing system are located on separate rigid circuitboards located in separate housing segments.
 6. The sensor of claim 4,wherein the analog impedance circuit and the digital processing systemare located on separate rigid circuit boards located in separate housingsegments.
 7. The sensor of claim 4, wherein the housing comprises threerigid housing segments, with the analog ECG circuit, the analogimpedance circuit, and the digital processing system each being locatedin one of the three housing segments.
 8. The sensor of claim 7, whereinthe digital processing system is located in a middle segment of saidthree housing segments, with each of the analog ECG circuit and theanalog impedance circuit being located in a respective outboard housingsegment and being connected to the digital processing system via arespective flexible circuit.
 9. The sensor of claim 2, wherein thesensing portion includes first and second pairs of electrode contactpoints, with 1) the first pair of electrode contact points including afirst voltage-sensing electrode contact point that is located in a firsthousing segment arranged to contact a first side of the patient's chest,and 2) the second pair of electrode contact points including a secondvoltage-sensing electrode contact point that is located in a secondhousing segment arranged to contact a second side of the patient's chestthat is laterally opposite to the first side of the patient's chest. 10.The sensor of claim 2, wherein the sensing portion includes first andsecond pairs of electrode contact points, with 1) the first pair ofelectrode contact points including a first current-injecting electrodecontact point that is located in a first housing segment arranged tocontact a first side of the patient's chest, and 2) the second pair ofelectrode contact points including a second current-injecting electrodecontact point that is located in a second housing segment arranged tocontact a second side of the patient's chest that is laterally oppositeto the first side of the patient's chest.
 11. The sensor of claim 2,wherein the sensing portion includes first and second pairs of electrodecontact points, with 1) the first pair of electrode contact pointsincluding a first voltage-sensing electrode contact point and a firstcurrent-injecting electrode contact point that are both located in afirst housing segment arranged to contact a first side of the patient'schest; and 2) the second pair of electrode contact points including asecond voltage-sensing electrode contact point and a secondcurrent-injecting electrode contact point that are both located in asecond housing segment arranged to contact a second side of thepatient's chest that is laterally opposite to the first side of thepatient's chest.
 12. The sensor of claim 1, wherein the analog ECGcircuit and the analog impedance circuit generate the ECG waveform andthe impedance waveform, respectively, using the same voltage sensed atthe voltage-sensing electrode contact point.
 13. The sensor of claim 11,wherein the analog ECG circuit and the analog impedance circuit generatethe ECG waveform and the impedance waveform, respectively, using thesame voltage sensed at the voltage-sensing electrode contact point, ofeach pair of electrode contact points, within each of the first andsecond housing segments.
 14. The sensor of claim 1, wherein the lightsource is configured to generate radiation in the red spectral rangenear 600 nm.
 15. The sensor of claim 1, wherein the light source isconfigured to generate radiation in the infrared spectral range near 800nm.
 16. The sensor of claim 1, wherein SYS is estimated from theequation: $\left. {SYS} \right.\sim\frac{m_{SYS}}{VTT}$ where m_(sys) isa patient-specific slope determined during a calibration that relatesSYS to VTT.
 17. The sensor of claim 1, wherein SV is estimated from theequation$\left. {SV} \right.\sim\frac{\left( \frac{{dZ}(t)}{dt} \right)_{\max}}{Z_{0}}$where (dZ(t)/dt)_(max) is the maximum value of a mathematical derivativeof the digital impedance waveform and Z₀ is the baseline value of thedigital impedance waveform.
 18. A sensor for measuring from a patient asystolic blood pressure (SYS), a diastolic blood pressure (DIA), and apulse oximetry (SpO2), the sensor configured to be located on thepatient's chest and comprising: a sensing portion having a flexiblehousing configured to be located on the patient's chest and enclosing abattery, a wireless transmitter, and all of the sensor's sensing andelectronic components, including: at least two pairs of electrodecontact points disposed on a bottom surface of the flexible housing,with each pair of electrode contact points comprising acurrent-injecting electrode contact point and a voltage-sensingelectrode contact point; an analog electrocardiogram (ECG) circuitcontained entirely within the flexible housing and in electrical contactwith a first voltage-sensing electrode contact point from a first pairof electrode contact points and a second voltage-sensing electrodecontact point from a second pair of electrode contact points, the analogECG circuit configured to generate an ECG waveform based on sensedvoltage; an analog impedance circuit contained entirely within theflexible housing and in electrical contact with current-injecting andvoltage-sensing electrode contact points in both the first and secondpair of electrode contact points, the analog impedance circuit beingconfigured to generate an impedance waveform based on sensed voltage; anoptical system located on a bottom surface of the flexible housing, theoptical system comprising a light source configured to generateradiation in both the red and infrared spectral ranges that separatelyirradiates a portion of the patient's chest disposed underneath theflexible housing, and a photodetector configured to detect radiation inthe red spectral range that reflects off the portion of the patient'schest to generate a first photoplethysmogram waveform (red-PPG), and todetect radiation in the infrared spectral range that reflects off theportion of the patient's chest to generate a second photoplethysmogramwaveform (infrared-PPG), with the optical system located between a firstelectrode contact point and a second electrode contact point of one ofthe two pairs of electrode contact points and configured to bereleasably connected to the patient's chest by the electrode contactpoints, such that the optical system irradiates the portion of thepatient's chest between the first electrode contact point and the secondelectrode contact point of one of the two pairs of electrode contactpoints; a first sensor for measuring stroke volume (SV) disposed withinthe flexible housing, the first sensor comprising a first processorconfigured to determine a value of SV from a maximum value of amathematical derivative of the impedance waveform and a baseline valueof the impedance waveform; a second sensor for measuring blood pressurecontained entirely within the flexible housing, the second sensorcomprising a processor configured to collectively processes theimpedance waveform and one of the red-PPG and infrared-PPG waveforms todetermine: 1) a first time point from one of the red-PPG andinfrared-PPG waveforms, 2) a second time point from the impedancewaveform, 3) a vascular transit time (VTT) from the temporal differencebetween the first and second time points, 4) a value of SYS from alinear equation that includes an inverse value of VTT, 5) a value ofpulse pressure (PP) from the product of SV and a calibration parameterrelating SV and PP, and 6) a value of DIA from the mathematicaldifference between SYS and PP; and a third sensor for measuring SpO2contained entirely within the flexible housing, the third sensorcomprising a processor configured to determine a value of SpO2 fromalternating and static components of both the digital red-PPG waveformand digital infrared-PPG waveform.